The present invention relates to an apparatus capable of determining the energy, position and time coordinates of light emission induced by interactions of gamma-rays in a planar array of discrete scintillator detectors having either a segmented or non-segmented light guide. The features of the present invention find particular application in the field of medical imaging whereby a single device can be used for Single Photon Imaging which includes traditional Gamma Cameras, Planar Imaging, Single Photon Emission Computed Tomography (SPECT) with or without Coincidence Photon Imaging and Positron Emission Tomography (PET). When operated in the SPECT mode, the present invention is comparable to existing high resolution SPECT systems. When operated in the PET mode, the present invention is an improvement over existing PET systems in that the device may be operated either in Pulse Height Discrimination mode or in Pulse Shape Discrimination mode thereby enabling depth of interaction encoding resulting in improved spatial resolution. Emission Computed Tomography (ECT) systems provide a means for sensing, and quantitatively measuring biochemical and/or physiological changes in the human body or other living organism. However, the use of the invention is not limited to such application.
Devices for detecting the distribution of gamma rays transmitted or emitted through objects to study the compositions or functions of the objects are well known to the art, e.g. the techniques referred to as Emission Computed Tomography can be divided into two specific classes; Single Photon Emission Computed Tomography (SPECT) uses radiotracers which emit gamma rays but do not emit positrons and Positron Emission Tomography (PET) which uses radiotracers that emit positrons. Therefore, the fundamental physical difference between the two techniques is that PET uses annihilation coincidence detection. The PET technique can determine, in-vivo, biochemical functions, on the injection of biochemical analog radiotracer molecules that emit positrons in a living body. The positrons annihilate with surrounding electrons in the subject body to produce a pair of gamma-rays, each having 511 keV of photon energy; traveling in nearly opposite directions. The detection of a pair of annihilation gamma-rays by two opposed detectors allows for the determination of the location and direction in space of a trajectory line defined by the opposite trajectories of the gamma-rays. Tomographic reconstruction is then used to superpose the numerous trajectory lines obtained by surveying the subject with an array of detectors to image the distribution of radiotracer molecules in the living body.
Emission Computed Tomography systems employ a variety of geometric configurations for the gamma-ray detectors. The choice of configuration is typically dictated by the manufacturer""s desired system performance and cost. The detector design must be capable of providing accurate estimates of gamma-ray energy, position coordinates, and in addition in the case of PET, coincidence time interval to reconstruct an image of the distribution of the radiotracer for in vivo studies. An example of such a device is disclosed in U.S. Pat. No. 4,750,972 to Casey et al., the disclosure of which is incorporated herein by reference and relied upon.
The position encoder and detector system disclosed by Casey et. al., is a two dimensional photon counting position encoder detector system, i.e., the array of scintillation crystals provides only the transverse coordinates of the photon interaction; the longitudinal photon interaction position of the excited scintillation crystal is undetermined. Photons impinging upon such detector systems at angles other than normal may traverse the path of several scintillation crystals resulting in uncertainty of their trajectory lines thereby degrading the image resolution due to parallax error.
A detector system capable of providing both the transverse and longitudinal position of photon interactions in scintillation crystals was disclosed in U.S. Pat. No. 4,843,245 by Lecomte. The approach involves the use of two scintillation crystals of different decay times which are stacked one upon the other. The position of photon interaction is determined by the Pulse Shape Discrimination technique. This method though capable of providing the transverse and longitudinal position coordinates of photon interactions in scintillation crystal detector systems will result in reduced system efficiency if the overall scintillator depth is constant for two different scintillator materials. If the scintillators are increased in length to compensate for the efficiency loss then the system resolution will be degraded.
Another approach to determine the transverse and longitudinal positions of photon interactions in scintillation crystal detector systems was disclosed in U.S. Pat. No. 5,122,667 by Thompson. The approach differs from that of Lecomte in that a single scintillator is used, further the method does not depend on decay time differences. The method employs the use of a scintillation light absorbing band located at the median interaction coordinate for a specific energy along the longitudinal axis of the scintillation crystal. The net effect is to divide the scintillation crystal into two regions whereby the photon is equally likely to interact. Pulse Height Discrimination is used to determine which of the two regions of the scintillator the photon interacted. This approach has the undesired effect of reducing the total collected scintillation light and of causing the Compton continuum of the high light yield scintillator to overlap the photopeak region of the low light yield scintillator. The result is inherent uncertainty in the contribution of scatter to the full energy photopeak.
In U.S. Pat. No. 5,349,191 Rogers discloses a method for determining the transverse and longitudinal position coordinates for interactions in scintillation crystal arrays which depends on the continuous variation of the total collected light with the longitudinal photon interaction coordinate of the light emission. The continuous variation in collected light requires a complex calibration of each detector as a function of longitudinal photon interaction coordinate from a collimated beam of photons directed at known positions along the length of the scintillator. This calibration method is difficult to implement for large arrays of scintillators.
In U.S. Provisional Application Ser. No. 60/037,519, filed on Feb. 10, 1997, and U.S. Provisional Application Ser. No. 60/042,002, filed on Apr. 16,1997, Moisan and Andreaco et. al. disclosed a device capable of determining the transverse and longitudinal coordinates of light emission induced by the interaction of photons in an array of photon detectors having a plurality of scintillation light guides. The device uses two or more layers of stacked scintillators all composed of the same scintillator material. Pulse Height Discrimination is used to determine which scintillator layer the photon interaction occurs. The device requires a difference in the light output from the two stacked scintillator layers of at least a factor of 1.5 times for the pulse height discrimination technique to be practicable. The approach has the undesired effect of causing the Compton continuum of the high light yield scintillator (which is nearest to the subject under study) to overlap the photopeak region of the low light yield scintillator. The result is inherent uncertainty in the contribution of scatter to the full energy photopeak.
The detector systems described in the above stated U.S. Patents when applied to medical imaging are specific to usage in PET. The predominant scintillator material is Bismuth Germanate (BGO), though other materials have been proposed or used (see Table 1). The SPECT detector systems are different in that Thallium doped Sodium-Iodide (NaI(Tl)) is used exclusively as the scintillator material. Further these systems use large continuous slabs of NaI(Tl) optically coupled to a continuous light guide. Anger logic is used for scintillation event localization. The exception to continuous NaI(Tl) slab detector systems for SPECT imaging was disclosed by Govaert in U.S. Pat. No. 4,267,452. This detector system is unique as a SPECT detector in that it is segmented. The segmentation of the NaI(Tl) is similar to PET block detector designs which use an active light guide. (For clarification detector light guides are of two general types: non-active light guides are composed of optical materials other than the scintillator; active light guides are composed of scintillator materials). The detector system disclosed by Govaert does not result in discrete scintillator elements whereby each element is a separate detector. Instead the segmentation process results in a block of NaI(Tl) that is subdivided into elements that share a common light guide of active scintillator material, i.e. the NaI(Tl) is not cut all the way through.
Other patents known to the inventors include the following:
The unique differences in SPECT and PET imaging modalities have resulted in detector designs which are suitable for their intended use in either SPECT or PET, but not both. However, the use of Fluorodeoxyglucose (FDG) with SPECT imaging systems has resulted in the application of SPECT detector designs in PET imaging. One problem in the application of SPECT detector designs in PET is that relatively thin scintillation crystals are preferred in Anger cameras to provide better intrinsic resolution and image detail. This results in poor detection efficiency in PET since the effective-Z and density of Nal(Tl) provides lower stopping power at 511 keV relative to PET scintillators (see Table 1). The efficiency of SPECT detector systems is further reduced by the use of absorptive collimation. The continuous slab of NaI(Tl) precludes the elimination of absorptive collimation.
SPECT detector system designs which are intended to bridge both SPECT and PET imaging modalities are known as hybrid devices. These systems have increased the NaI(Tl) scintillator thickness for higher efficiency and have added coincidence detection circuitry and attenuation corrections. Despite these changes the continuous slab of Nal(Tl) scintillator detector designs are inferior to PET specific detector designs in terms of system performance.
The hybrid SPECT detector designs have compromised their SPECT performance while providing inferior PET performance. A need has arisen for a hybrid PET/SPECT detector system which provides state of the art SPECT and PET system performance which does not suffer from the heretofore stated disadvantages.
Accordingly, it is an object of the present invention to provide a detector system design which does not suffer from the heretofore stated disadvantages.
Another object of this invention is to provide a detector system design for the detection of radioactive events inclusive of single photon emitters and positron emitters.
A further object of this invention is to provide a detector system for imaging radioactive event distributions.
A still further object of this invention is to provide a detector system for planar and tomographic imaging.
It is yet another object of the present invention to provide a detector system for use in both SPECT and PET imaging.
A further object of the present invention is to provide a detector system with depth of interaction encoding.
It is another object of the present invention to provide a detector system with active shielding against background and scatter radiation.
A further object of the present invention is to effectively eliminate scintillator self-radiation by the use of pulse shape discrimination.
A still further object of the present invention is to provide a detector with time-of-flight encoding in one or more embodiments of the design.
Other objects and advantages of the present invention will become more apparent upon review of the detailed description and associated drawings of the scintillator detector array for encoding the energy, position and time coordinates of gamma-ray interactions.
In accordance with the various features of this invention, a scintillation detector is provided which includes a plurality of discrete scintillators composed of one or more scintillator materials. The discrete scintillators interact with incident radiation to produce a quantifiable number of photons with characteristic emission wavelength and decay time. A light guide is operatively associated with the scintillation crystals and may be either active or non-active and segmented or non-segmented depending upon the embodiment of the design. Photodetectors are provided to sense and quantify the scintillation light emissions. The process and system embodying various features of the present invention can be utilized in various applications such as SPECT and PET imaging systems. In accordance with the present invention, the detector array of the present invention incorporates either a single layer of discrete scintillators or discrete scintillators composed of two stacked different layers that can be the same scintillator material or of two different scintillator materials. In either case the different layers are composed of materials that have distinctly different decay times. The variants in these figures are the types of optical detectors which are used, i.e. photomultipliers and/or photodiodes, whether or not a segmented optical light guide is used, and whether the light guide is active or non-active. If a segmented optical light guide is used then the variant is whether the configuration is inverted or non-inverted.